Rapid acquisition magnetic resonance imaging using radial projections

ABSTRACT

High resolution and high speed MR imaging is provided for imaged objects in which the brightness of the imaged objects dominates the surrounding tissues by using sparse angular sampling and projection acquisition techniques. Individual objects throughout a large field of view are imaged at a rate and resolution normally associated with small field of view techniques. For applications such as angiography, artifacts associated with sparse angular sampling are acceptable. Volume images are acquired by combining sparsely sampled projection in two dimensions with weighted Fourier Acquisition in the third dimension.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application claims the benefit of U.S. provisional application No.60/069,430 filed Dec. 12, 1997 and U.S. provisional application No.60/081,409 filed Apr. 10, 1998 both hereby incorporated by reference.

STATEMENT REGARDING FEDERALLY SPONSORED RESEARCH OR DEVELOPMENT

“This invention was made with United States Government support awardedby the National Institute of Health under grant NIH HL 52747. The UnitedStates has certain rights in this invention.”

BACKGROUND OF THE INVENTION

The field of the invention is magnetic resonance imaging (“MRI”) methodsand systems and more particularly a method and apparatus for rapidlyacquiring MRI data from a portion of an imaged object.

MRI Imaging

When a substance such as human tissue is subjected to a uniform magneticfield (polarizing field B₀) the individual nuclei in the tissue attemptto align their magnetic moments with this polarizing field but as aresult of nuclear spin, precess about it in random order at theircharacteristic Larmor frequency. The Larmor frequency is dependent onthe strength of the magnetic field and on the properties of a particularnucleus as represented by a magnetogyric constant γ. Nuclei whichexhibit this phenomenon are referred to as “spins”.

By convention, the polarizing field B₀ is considered to lie along a zaxis of a Cartesian coordinate system. The procession of the nuclei inthe polarizing field B₀ creates a net magnetic moment M_(z) in thedirection of the polarizing field. Individual spins have magneticmoments that are perpendicular to the z axis in the transverse or x-yplane, however, the random orientation of the spins cancels any nettransverse magnetic moment.

In MRI imaging, a radio frequency signal is applied in the x-y planenear the Larmor frequency to tip the net magnetic moment into the x-yplane so that it rotates at the Larmor frequency. The practical value ofthis phenomenon resides in the signal which is then emitted by theexcited spins termed the NMR signal (“nuclear magnetic resonance”). Insimple systems, the excited spins induce an oscillating sine wave in areceiving coil which may be the same coil used to excite the spins. Theamplitude of this signal decays as a function of the homogeneity of themagnetic field caused by atomic scale interaction between the spins or“spin-spin” relaxation and the engineering limitations of producing atruly homogenous polarizing field B₀. This decay is caused by a loss ofphase coherence in the spins and is commonly referred to as T*₂relaxation. Second decay mechanism is the gradual return of the magneticmoments of the individual spins to a longitudinal direction aligned withthe polarizing field B₀. This is termed T₁ relaxation and in mostsubstances of medical interest is much longer than T₂ relaxation.

An image of a patient may be obtained by evaluating the NMR signalcontributed by different spins at different locations in the patient'stissue. A pulse sequence using gradient magnetic fields encodes locationinformation on the spins in the form of the phase and frequency. Theencoded spin signal may then be separated to produce an image.

A wide variety of pulse sequences is known. For example, the spin warpor spin echo technique is described in “Spin Warp NMR Imaging AndApplications To Human Whole-Body Imaging” by W. A. Edelstein et al.,Physics in Medicine and Biology, vol. 25, pp. 751-756 (1980); the steadystate free precession (“SSFP”) technique including gradient refocusedacquired steady state pulse sequences (“GRASS”) as described in U.S.Pat. No. 4,665,365 and contrast enhanced fast imaging (SSP-ECHO)described in “Rapid Fourier Imaging Using Steady State Free Precision”,R. C. Hawks and S. Patz, Magnetic Resonance in Medicine 4, pp. 9-23(1987); and echo planer imaging (“EPI”) is described in an article byPeter Mansfield (J. Phys. C. 10: L55-L58, 1977). These descriptions ofpulse sequences are hereby incorporated by reference.

Cartesian MRI Acquisition

In a representative spin echo pulse sequence, a z-axis gradient and anarrow-band, radio frequency excitation pulse may be applied to apatient, for example, so that only spins in a “slice” to the patientperpendicular to the z-axis are excited. An x-gradient field may then beapplied to cause the spins at one side of the slice to precess fasterthan spins on the other side of the slice. In this manner, the spinshave been given a frequency encoding that allows them to bedistinguished along the x-axis.

This NMR signal of the spins at different frequencies is acquired for aperiod of time and digitized to provide a first row of data that may bestored in an array in a reconstruction computer. The number ofdimensions of the array and the number of elements in the array define ak-space well known to those in the art. The NMR signals must be sampledat a rate at least twice the frequency of the highest frequencycomponent of the NMR signal (the Nyquist sampling rate) so as to preventthe introduction of aliasing artifacts.

Additional NMR signals are then collected for this slice with the samex-gradient but with a progressively increased y-axis gradient field. They-axis gradient serves to phase encode the spins in the y-direction.Each successive NMR acquisition with a different y-axis gradient forms asuccessive row in the k-space array in the computer.

Once k-space has been filled, a two dimensional Fourier transform may bemade of the k-space data to produce the desired image. Generally it isknown to band limit the NMR signal to eliminate the contribution ofspins beyond certain spatial ranges in the frequency encoding x-axisdirection limiting the amount of k-space data somewhat. Such bandlimiting cannot be performed in the phase encoding direction, however,until after the data is fully acquired and therefore is of little valuein reducing the data acquisition time.

Radial MRI Acquisition

In an alternative method of data acquisition, the k-space data is fillednot by rows and columns but by a series of radial projections about apoint within k-space. This acquisition technique is analogous to theacquisition of data in an x-ray computed tomography (“CT”) machine andallows the data to be reconstructed into an image by CT-type algorithmsincluding filter back projection.

MRI Angiography

In MRI angiography, images of the blood vessels are obtained. Forcontrast-enhanced applications in which contrast materials such asgadolinium compounds are injected into a peripheral vein, theacquisition of k-space data must be carefully coordinated with thearrival of contrast so as to prevent an unfavorable variation in theweighting of the k-space data. The availability of a high speed imagingtechnique would be helpful in this regard since it would permit a seriesof images to be obtained throughout the passage of contrast.

In contrast enhanced MRI, two images, one before the introduction of acontrast medium into the vessels and one after the introduction of thecontrast medium, may be obtained and subtracted. The subtracted imagereveals information about the bloodflow through the vessels allowing thedetection of obstructions and the like. Structures other than flowingblood are similar in the two images and thus substantially reduced incontrast.

The timing of the acquisitions of the two MRI images is crucial toproviding a high contrast image. Normally there is a time delay betweenthe introduction of the contrast medium into the patient and its time ofarrival at the region of the vessel of interest.

Ideally the first image should be concluded immediately before thearrival of the contrast medium so as to provide an accurate comparisonimage and the second image begun immediately after the arrival of thecontrast medium so as to be complete before the contrast mediumdissipates The time consuming process of acquiring an image of a patientand the difficulty of monitoring the progress of the contrast medium,make production of a high quality contrast enhanced MRI image adifficult task.

Acquisition Speed

It would be advantageous to be able to acquire images of higherresolution more quickly. This is important when the available imagingtime is limited by the passage of injected contrast material or byrespiratory motion. In Cartesian acquisitions image spatial resolutionis proportional to imaging time, so any reduction in acquisition timeproduces images of reduced spatial resolution. Within the context ofCartesian imaging, some investigators have developed methods to imagereduced field of view. Reduction of the field of view reduces the amountof data required. Therefore, imaging time can be reduced. A disadvantageof this is that in Cartesian imaging objects from outside the field ofview can appear inside the selected small field of view due to aliasing.In Cartesian imaging this aliasing results in an artifact which looksexactly like the object from outside the field of view. Although bandlimiting can be used to reject objects from outside the field of view inthe frequency direction, this is not possible in the phase encodingdirection because each row of phase encoding data relates to spinssituated throughout the field of view.

One method of contending with this problem is the method of Hu andParrish (X. Hu, T. Parrish, Reduction of FOV for dynamic imaging. Magn.Reson. Med. 31, 691 (1994). In this method, a preliminary image of theentire field of view is obtained. The portion of this imagecorresponding to the desired field of view is set to zero. Then theremaining data are subjected to a Fourier transformation to provide ak-space data set corresponding to just that material which might aliasinto the small field of view. During dynamic imaging of the small fieldof view, this k-space data set is subtracted from the data associatedwith the small field of view. This removes aliased signals from thesmall field of view, but only if the objects outside the small field ofview are truly stationary.

In projection MRI acquisitions there is no phase encoding direction, assuch, and it has therefore been proposed to use such an acquisition tolimit the field of view thereby reducing acquisition time. Eachprojection is bandlimited to limit the inclusion of spins from outsidethe region of interest along the projection, however spins in regionsperpendicular to the axis of projection cannot be so eliminated andproduce image artifacts. These artifacts must be addressed by estimationtechniques or supplementary measurements of the out of region areas forlater cancellation. Such a technique is described in “Zooming by BackProjection” by K. Scheffler in the Proceedings of ISMRM, FifthScientific Meeting and Exhibition, Vancouver. BC Canada, Volume 1 page288. Scheffler and Hennig (K. Scheffler, J. Hennig, Reduced circularfield-of-view imaging. Magn. Reson. Med. 40, 474-480 (1998).) haveapplied the Hu/Parrish algorithm using projection acquisition. In thiscase, as in the Cartesian case, the imaging speed for a given resolutionis increased. However, dynamic changes can only be viewed within thereduced field of view, and aliasing occurs if the outer material is notstationary. These techniques are limited by the fact that only a smallfield of view is acquired.

In projection imaging, the requirement that the angular samplinginterval be equal to the radial sampling interval is that the number ofprojections NP is related to the number of radial samples NR by

NP=NR·π/2

This requirement poses an inherent time disadvantage for projectionimaging relative to Cartesian imaging. The fact that it requires anadditional factor of π/2 to acquire an image of the same resolution isone of the reasons that Cartesian acquisition is the primary method formagnetic resonance imaging in spite of the fact that the first MRIimages in the mid 1970s were acquired with projection acquisition.

It is well known in the X-ray computed tomography literature, whereprojection acquisition is used, that acquisition time can be increasedif sparse angular sampling is used. In this method the number ofprojections is decreased. It is also well known that as the number ofangular samples is decreased, spatial resolution is not decreased butradial streak artifacts emanate from all objects within the field ofview. In X-ray CT where bone provides a dominant signal which farexceeds the tissue signal differences to be distinguished, the presenceof such artifacts is completely unacceptable. Up until this point intime, it has been assumed by the MRI community that such artifacts wouldbe similarly unacceptable in magnetic resonance imaging if sparseangular sampling were to be used. Therefore, the small field of viewtechniques mentioned above have been resorted to in order to increaseimaging speed in the limited number of situations in which the field ofview can be compromised.

BRIEF SUMMARY OF THE INVENTION

The present inventors have recognized that when the NMR data is acquiredin projection rather than Cartesian fashion the rate at which spatialresolution can be acquired is significantly increased. As can beinferred from X-ray computed tomography, spatial resolution iscompletely determined by the resolution in the readout direction withineach projection and not by the number of acquired projections. As thenumber of projections is decreased, resolution is unaffected. The onlyeffect is an increase in artifacts The artifacts generated by the sparsesampling in projection imaging are different from those generated inCartesian acquisitions and can be readily tolerated in a number ofimportant imaging applications.

Around each object within the overall field of view there is a smallregion (local field of view) in which the object does not produce anyartifacts. The size of this artifact-free region does depend on thenumber of acquired projections. Outside of this small region each objectdoes produce streak artifacts which can enter the local fields of viewassociated with other objects. The present inventors have recognizedthat these artifacts are typically no more than a few percent of thesignal associated with the object producing them and that forapplications such as angiography, pancreatography or bile duct imaging,where the signals of interest are the most dominant signals in theoverall field of view, these artifacts appear to be completelytolerable. In such situations azimuthally undersampled projectionimaging provides the speed and resolution advantages of reduced field ofview imaging, but does so simultaneously throughout an entire largefield of view. Preliminary results suggest that a speed increase of afactor of six is often possible.

Specifically, then, the present invention provides a method of MRIimaging of structures, the structures providing NMR signals withintensities dominating the intensity of other NMR signals of othermaterials within a field of view. The method includes the steps ofgenerating a series of gradient fields along axes distributed over arange of angles about an axis in the field of view, then acquiring NMRsignals of the field of view at different gradient fields, eachacquisition providing an angular projection of data in k-space havingdata points radially spaced at distances and are along a projection. Thenumber of acquisitions is limited to a number of projections less thanNRπ/4 in number (where NR is the number of radial samples) so thatk-space is sparsely sampled. The projections are reconstructed toproduce a field of the image of view displaying the entire field ofview.

It is thus one object of the invention to provide a rapid NMRacquisition technique well suited to bright objects dominating otherobjects in the field of view. Such situations may include imaging bloodvessels in angiography, or imaging the pancreas or liver ducts. Thepresent invention provides a dynamic view of the entire field of view,not just the object.

The projection acquisition may be combined with a volumetric acquisitionin which a phase encoding in gradient is applied along the axis duringthe acquisition of the NMR signals and a Fourier reconstruction is madealong that axis to provide a volume image.

Thus it is another object of the invention to provide for the benefitsof projection acquisition together with the benefits resulting fromweighted k-space acquisition along the phase encoding gradient. As isknown in the art, it is desirable to increase k-space sampling for lowerk-space frequencies at the expense of higher k-space frequencies.Projection imaging does not allow such k-space weighting in the slice ofthe projection since each projection acquires data near the center ofk-space. However, when combined with Fourier reconstruction in thevolumetric axis, such weighting can be obtained. The phase encoding maybe performed after a full set of projection images are obtained or maybe performed in between each projection image. In this latter case, theprojection angle selected may alternate between two or more interleavedsets so as to minimize the time required to obtain a sparse sampling ofk-space yet to allow the more complete sampling as time permits.

In the case where the structure being imaged is smaller than acontaining imaged object, the step of spatial saturation of two bandsparallel to the gradient direction and separated by the interveningwidth of the structure may be included.

It is therefore another object of the invention to provide a method ofsuppressing the artifacts produced by sparse sampling with projectionimaging.

During the acquisition of the NMR signals, a contrast medium may beintroduced into a field of view and used to produce a contrast indexindicating the arrival of a contrast medium in the structure. Thiscontrast index may be displayed and/or used to select particular NMRsignals for the generation of low and high contrast images. The low andhigh contrast images may be subtracted, for example, to performsubtraction angiography.

Thus it is another object of the invention to provide a real-timemeasure of contrast. The NMR signals acquired at each projection may besimply integrated to produce a contrast indication. The contrastindication may be weighted according to the particular angle so as toremove anatomical effects or when a comparison value is being produced,the comparison may be made only between corresponding angles.

The foregoing and other objects and advantages of the invention willappear from the following description. In this description, reference ismade to the accompanying drawings which form a part hereof and in whichthere is shown by way of illustration a preferred embodiment of theinvention. Such embodiment does not necessarily represent the full scopeof the invention, however, and reference must be made therefore to theclaims for interpreting the scope of the invention.

BRIEF DESCRIPTION OF THE SEVERAL VIEWS OF THE DRAWINGS

FIG. 1 is a simplified representation of an MRI scanner showing apolarizing magnet and three stationary orthogonal gradient coils alignedwith the z-axis of a Cartesian coordinate system and attached togradient amplifiers, the figure further showing RF transmission andreceiving processing circuitry of the present invention controlled byand communicating with a processor having a display and a manual inputdevice for setting a location of a region of interest;

FIG. 2 is a perspective graphical representation of the polarizingmagnetic field with respect to three orthogonal gradient fields asproduced by the MRI scanner of FIG. 1;

FIG. 3 is a schematic representation of an x-y plane through the bore ofthe magnet of FIG. 1 showing the generation of an arbitrarily angledgradient field along axis x′ by means of a linear combination of thefixed x- and y-axis gradient fields;

FIG. 4 is a figure similar to FIG. 3 showing multiple, angled gradientaxes extending radially through a region of interest displaced from amagnet center;

FIG. 5 is a figure similar to FIG. 3 depicting regions of frequencyencoding along an axis x′ and various dimensions relevant to theacquisition of NMR data from that field of view;

FIG. 6 is a flowchart describing steps executed by the processor of FIG.1 in acquiring an image per the present invention;

FIG. 7 is a representation of NMR data obtained in a prior art MRI scanarranged in k-space and as stored in an array in memory of the processorof FIG. 1;

FIG. 8 is a figure similar to FIG. 7 showing NMR data acquired per thepresent invention as arranged in k-space;

FIGS. 9, 10 and 11 are line graphs along lines through a reconstructedimage from the data of FIG. 8 having respectively, an added bowlartifact, the bowl artifact in isolation, and the image data with thebowl artifact subtracted from it;

FIG. 12 is a figure similar to that of FIG. 6 showing an alternativeembodiment using spin suppression to reduce image artifacts and/or imagesubtraction;

FIG. 13 is a figure similar to that of FIG. 4 showing regions of spinsuppression and the Fourier transform of a spin saturation RF pulse usedfor spin suppression;

FIG. 14 is a graphical representation of a gradient sequence suitablefor the spin suppression of FIG. 13 and the acquisition of NMR data;

FIG. 15 is a schematic representation of a series of acquisitions perthe gradient sequence of FIG. 14;

FIG. 16 is a combination time graph of contrast produced by an injectedcontrast medium and of NMR data acquisition in subsets according to thepresent invention; and a schematic representation of the collection ofthe subsets into subtractable images;

FIG. 17 is a schematic representation of the division of a projectionset of NMR data into the subsets of FIG. 16; and

FIG. 18 is a perspective representation of a volume of acquired data ink-space showing the division of the volume into kz-axis regions.

DETAILED DESCRIPTION OF THE INVENTION MRI System Components

An MRI system 10 of a type suitable for use with the present inventionmay include an annular superconducting magnet 12 providing a homogenousB₀ magnetic field along a z-axis within a central bore 14 of the magnet12. The bore 14 may generally define a maximum field of view 62 of theMRI instrument.

Attached to a coaxial form 16 positioned within the bore 14 are x, y-andz-gradient coils 18, 20 and 22, respectively. These coils producegradient fields also directed along the z-axis but having spatialgradation measured along the x, y-and z-axis respectively. An RFtransmitting and receiving coil 24 is also positioned coaxially withinthe bore 14 to allow for excitation of the nuclear spins and a detectionof the resulting NMR signal. The construction of magnet and coilstructures as described is well known in the art and is intended to beonly representative of a broad class of instruments.

Each of the gradient coils in x, y, and z is connected to acorresponding gradient driver circuit 26 that may receive commands froma central processor 29 so as to provide an electrical signal controllingthe gradient produced by the coils 18, 20 and 22 both in time andamplitude. The gradient coils are fixed in relationship to the bore sothat the magnetic gradient fields produced remain symmetrical about abore axis 28, but the slope of the gradients is controlled by thegradient driver circuit 26. Thus, referring both to FIGS. 1 and 2, thesuperconducting polarizing magnet 12 produces a B₀ field 30 uniformacross the x, y and z-axes within the field of view. A y-gradient 32provides a magnetic field whose vectors are also aligned with the z-axisbut which are directed counter to the B₀ field 30 at low y-axis values,with the B_(o) field at high y-axis values, and which vary linearlybetween the two y-axis extremes. Likewise the x-gradient 34 andz-gradient 36 vary linearly between opposition and alignment with the B₀field as one moves to higher x-and z-values. The zero point of thegradients is generally aligned with the bore axis 28 of the magnet 12.

Referring again to FIG. 1, also connected with processor 29 is an RFexcitation circuit 38 which receives a control signal from the processor29 to produce an RF excitation pulse to the RF coil 24 to excite spinsof a patient or the like in bore 14 into resonance as has been describedand as is generally understood in the art. After completion of the RFexcitation pulse, one or more NMR signals may be acquired as received bythe coil 24 through RF amplifier 40 as is understood in the art. The RFsignal so received is processed by a programmable frequency offsetcircuit 42 which provides a heterodyning action to scale downward thefrequencies in the RF signal by an offset amount 44 received from theprocessor 29. The NMR signals are next filtered by a programmablelow-pass filter 46 which receives its filter break point 48 from theprocessor 29. The resulting NMR signal is then sampled and digitized andstored in memory in the processor 29 in an array as will be describedbelow.

It will be understood that the programmable frequency offset circuit 42and the programmable low-pass filter 46 may be realized as discreteelectronic circuits or portions or all of these elements maybe realizedin software as operations performed on the digitized and sampled NMRdata according to techniques well know in the art. Likewise, thesampling rate may be realized after data acquisition at a highersampling rate by an averaging of multiple samples into one or othersimilar techniques. Because the invention may be realized in software,no change in the hardware of the MRI scanner is needed. The order of theprocessing of the NMR signal by the programmable frequency offsetcircuit 42 and the programmable low-pass filter 46 can be reversed, butis preferably as shown for reasons that will become apparent.

The processor 29 receives the NMR data and reconstructs it into an imagethat may be displayed on a display 50. The processor also receives,among other control and console commands not described herein, aposition signal 52 from a cursor control device 54 such as a mouse,track ball or joystick.

Reduced Field Projection Acquisition

While the present invention is not limited to imaging a reduced field ofview and in fact provides a dynamic imaging of the entire field of view,the inventors have provided certain advances applicable also to reducedfield of view imaging as will now to be described.

Referring now to FIGS. 3 and 6, the processor 29 executes a storedprogram 58 contained in memory in order to acquire NMR data according tothe present invention. In a first step of this program 58, indicated byprocess block 60, a region of interest 64 is identified within themaximum field of view 62 of the instrument. The maximum field of view 62will generally be the entire portion of the bore 14 in which a patientor other imageable material is placed, whereas a region of interest 64may represent a portion of a patient, for example a part of the colon,substantially smaller in area than the maximum field of view 62. Theregion of interest 64 may be offset from the bore axis 28 of the bore 14in the x-and y-direction by amount x₀ and y₀. The region of interest 64may be selected by a physician simply by knowledge of the coordinates ofinterest or interactively with the physician operating the cursorcontrol device 54 to move a cursor 66 or the like on the display 50where the display shows a previously obtained slice image of the patientor a similar schematic representation. The area and shape of the regionof interest 64 may also be input by the physician. Typically, and asassumed herein, the region of interest 64 will be circular but it neednot be.

Referring again to FIG. 6, after the region of interest 64 is identifiedboth as to location and dimension, calculations for a set of radialgradient fields and associated offset frequencies and/or filter settingsare generated as indicated by process block 63. During the acquisitionof the image data, the radial gradients will be applied to the maximumfield of view 62, including the region of interest 64 along axes x′ atangles θ with respect to the x-axis. In the preferred embodiment, theradial gradients will be distributed over 180° at equal angular spacingsand of a number determined by the size of the region of interest.

Each radial gradient may be generated by energizing at predetermineddriving amplitude, one or more of the fixed gradient coils 22, 20 and18. Thus referring to FIG. 3, a gradient along x′ at an arbitrary angleθ may be produced by simultaneously energizing gradients x and y suchthat their relevant amplitudes G_(x) and G_(y) are according to thefollowing equation:

G _(y) =G _(ymax) sin(θ)

G _(x) =G _(xmax) cos(θ)

where G_(ymax) and G_(xmax) are the maximum values of G_(y) and G_(x),respectively. Clearly gradients x′ tipping out of the x-y plane may alsobe obtained by energizing the z-gradient for the acquisition of data inother planes than the x-y plane or in a volume rather than a slice.

Referring now to FIG. 4, gradients x′ distributed at angles about thebore axis 28 will also provide gradients oriented at these same anglesabout the center of the region of interest 64, albeit the zero point ofthe gradient fields will not be centered on the region of interest 64.Rather, referring to FIG. 5, for a given gradient along an axis x′, theisocenter or zero point of the gradient (at bore axis 28) will differ byan amount 68 from the center of the region of interest 64 where theamount 68 causes a frequency offset in the acquired NMR signal as afunction of θ as follows:

Frequency offset(x ₀ , y ₀)=γ(G _(max)(y ₀ sin θ+x ₀ cos θ))

where G_(max)=G_(ymax)=G_(xmax)

Thus, the spins centered in the region of interest 64 will have adifferent frequency than the spins at the bore axis 28 depending on theangle θ at which the gradient x′ is oriented. For this reason, the NMRsignals acquired will be adjusted in frequency by this offset amount(that varies with θ) so that they may be properly combined forreconstruction.

Truncation of NMR Data Along x′

Referring again to FIG. 5, an NMR signal acquired with a gradient alongaxis x′ will be frequency encoded by the composite x′ gradient field.Thus, before the frequency shifting described above, spins for lower x′values 70 outside of the region of interest 64 will have a lowerfrequency and the spins for higher x′ values 72 also outside of theregion of interest 64, will have a higher frequency. After the frequencyshifting described above, the lowest frequency spins will be in theregion of interest 64 and the spin frequency in the lower x′ values 70and the higher x′ values 72 will be equal and opposite in phase andhigher in magnitude than the spins in the region of interest 64. Thusthese spins may be removed with a low pass filter whose break-point isdetermined by the size of the region of interest 64. Prior toreconstruction or even sampling of the data (if a discrete rather thansoftware filter is used), contributions from spins outside the region ofinterest 64 along the x′ axis may be eliminated. The effectivebandlimiting of the NMR data allows a lower effective sampling ratewithout aliasing and permits an improved signal to noise ratio in thesampled points 83.

Referring again to FIG. 6, the offset frequencies and filter breakpointneeded for each gradient at x′ are thus computed at process block 63prior to the generation of the gradients and acquisition of NMR data sothat the acquisition of NMR data is not slowed by these calculations. Itwill be understood, however, that these computations may also be doneduring the acquisition of NMR data with suitably fast hardware.

Once the necessary offsets and breakpoints are determined, NMR data isacquired at each gradient angle θ and then shifted by the offsetfrequency by programmable frequency offset circuit 42 and filtered bythe programmable low-pass filter 46, per the parameters previouslydetermined as indicated by process block 74 and the loop of decisionblock 76. The acquisition is complete once NMR data has been obtainedalong each gradient axis x′.

Referring now to FIG. 8, the data of each NMR signal 78 is stored in anarray 80 in memory of the processor 29 whose dimensions conform tok-space as is understood to those of ordinary skill in the art.Generally k-space differs from the Cartesian space of the bore 14 inthat it is a spatial frequency domain rather than a time or positiondomain. The array 80 has rows of constant ky dimension and columnsextending across the kx dimension. As each NMR signal 78 is acquired, itis stored in the array 80 at an angle from the kx axis corresponding tothe angle that the gradient x′ is to the x-axis. Thus, data fills thearray 80 as a series of spoke-like projections 82 angled in k-spacecrossing rows and columns.

Referring momentarily to FIG. 7, this radial acquisition is in contrastto a more typical filling of k-space on a row by row basis, each rowrepresenting the NMR signal 78, sampled at points 83, taken for adifferent phase encoding provided by the y-gradient.

Comparing these methods we note that there is a significant differenceregarding the opportunity to exclude signals from unwanted regions. Thishas important consequences regarding signal to noise ratio, scan time orspatial resolution. When it is desired to obtain signals from a reducedfield of view, it is necessary to worry about the extent to whichsignals associated with regions outside the field of view can alias intothe desired field of view. In the conventional case employingrectilinear readout, temporal frequency filtering may be applied to thereal-time readout signal to exclude signals outside a desired range ofx-values, since x position is uniquely related to temporal frequency.However, since the y gradients are applied intermittently and are offduring readout, there is no relationship between signal temporalfrequency and y position. Therefore aliasing can occur in the ydirection. In the case of projection reconstruction, the x′ axis rotatesand signals outside a given x′-range may be excluded. When the rotatingsaturation regions (described later in FIG. 13) are applied, aliasingfrom signals outside the desired field of view but within the limitedspatial region(limited by the x′ temporal filter) extending in the y′direction are also excluded. When small fields of view are successfullyisolated, the k space step size may be increased (field of view=1/deltak). This means that fewer k-space steps are required to reach the edgesof k-space. Because of this, readout of the signal using lower bandwidthcan be employed to reduce noise. Therefore the more successful isolationof the small field of view using the described projection methods can beused to reduce noise. Optionally, if bandwidth reduction is notexploited, the same bandwidth can be used to reach the edges of k-spacefaster, resulting in reduced scan times, or for equal scan times,additional k-space data can be recorded, resulting in higher spatialresolution.

Referring again to FIG. 6, once a desired number of projections 82 havebeen acquired, the data may be reconstructed into an image by a numberof techniques well known in the art of computed tomography (“CT”)imaging. In the preferred embodiment, the reconstruction technique is“filtered back-projection” where a filter is applied to each projection82 of data and it is back projected or “smeared” across the image planeto develop an image, a technique which is well known in the art.Alternatively, Fourier reconstruction techniques may be used in whichpoints filling k-space are interpolated between the points 83 of theprojections 82 and a two-dimensional Fourier transform performed. Thereconstruction of the image is indicated at process block 86. The imageis displayed on display 50 and may be superimposed on a template showingthe larger field of view as has been described.

Spins Outside The Field Of View

Although frequency filtering of the NMR signal eliminates thecontribution to the acquired NMR signals of spins outside the region ofinterest 64 along the x′ axis this is not true for spins displaced alongthe y′ axis (perpendicular to the x′ axis) outside the region ofinterest 64. These spins produce a bowl artifact in the reconstructedimage. Referring to FIGS. 6 and 9, in a first embodiment, at nextprocess block 88, a bowl artifact may be subtracted out from the image.The bowl artifact is a radially symmetric, generally upwardly concaveoffset centered in the image of the region of interest 64. The exactfunction of a bowl function will vary with the filter used forreconstruction and certain characteristics of the machine and imagetissue but in a preferred embodiment is taken to be the followingequation:

I _(bowl)(r)=I _(ring)(R−r+1)^(−0.68)+(R+r−1)^(−0.68)

where

I_(bowl) is the bowl artifact;

I_(ring) is the intensity of the image at the edge of the reduced fieldof view;

R is the radius of the field of view; and

r is the distance from the center of the region of interest.

By subtracting the bowl artifact 92 from the reconstructed image 90, apure image 94 is obtained which may then be displayed as indicated byprocess block 96.

Spin Suppression

Referring now to FIGS. 12 and 13, in an alternative embodiment, the bowlimage artifact is reduced by the suppression of spins outside the regionof interest 64. At process block 74′ corresponding generally to processblock 74 previously described, a saturation pulse sequence shown in FIG.14 and to be described, saturates and dephases spins of the maximumfield of view 62 for values along the y′ axis outside the region ofinterest 64. These saturated and dephased regions 67 are as indicated bythe cross hatching and rotate with the rotation of the x′ axis duringacquisition. Through this technique, spins outside of the region ofinterest 64 that are aliased into the image of the region of interest 64are reduced. The combination of saturated and dephased region 67 and thefrequency filtering which eliminates the contribution of spins fromregions 65 along the x′ axis outside region of interest 64,substantially eliminates such aliasing.

Referring to FIG. 14, one example of a pulse sequence for accomplishingthe spins suppression described above excites the spins in a singlez-axis slice using RF pulse 100 during the application of z-gradient 102used to select the slice. After x and y gradient prewinder pulses and az gradient rewinder pulse 103, a combination of x and y gradients 104may be applied to generate the x′ axis as previously described alongwhich an NMR signal 106 will be obtained. The NMR signal 106 provides afirst projection 108 as shown in FIG. 15 and the data of a firstprojection 82 shown in FIG. 8.

Generally, as shown in FIGS. 14 and 15, after one or more projections108 are obtained, a suppression pulse sequence 110 is performed. In thesuppression pulse sequence 110, the x and y-gradient fields are switchedto values 107 to produce a gradient along the y′ axis generallyperpendicular to the next anticipated x′ axes at which NMR signals 106will be obtained. Referring also to FIG. 13, at this time a Hadamard RFpulse 109 is generated to excite spins only in the dephased regions 67so that their moments are precessing in the x-y plane. The Hadamardpulse 109 is a combination of frequencies p(t) which in combination withthe gradient along y′ do not excite the spins within the region ofinterest 64 along the y′ axis as will be understood to those of ordinaryskill in the art. Following the Hadamard RF pulse 109, the x andy-gradients are set to arbitrary values to cause a dephasing of thespins in dephased regions 67.

A suppression pulse sequence 110 need not be performed after everyacquisition of data as only a very small wedge of unsuppressed spinsoutside the region of interest 64 will occur with incremental changes inthe advancing of the x′ axis, and this wedge of spins may be ignored infavor of faster data acquisition.

Entire Field Projection Acquisition

As mentioned, the present invention is preferably not limited to imaginga reduced field of view but instead provides a dynamic imaging of theentire field of view. Referring again to FIG. 6, in the presentinvention, blocks 60 63, and 80 may be eliminated and a set ofacquisitions may be acquired along projections per process block 74 withbandlimiting only to the extent normally done in Cartesian acquisitionsso as to admit spins from the entire field of view. The number ofprojections is limited to a number that provides a sparse sampling ink-space. Referring to FIG. 8, the present invention contemplates thatthe number of projections 82 is limited to an amount in number less thanNRπ/4 where NR is equal to the number of samples in each projection.Artifact free projection imaging would normally be expected to use NRπ/2projections. The present inventors have recognized that this sparsesampling does not decrease the resolution of the image appreciably butsimply creates image artifacts. As a result, there is no need to limitthe field of view because the entire field of view can be imaged atessentially no additional cost in time. The invention thus provides adynamic view of the imaged object outside the region of interest such asmay be important of the physician.

Further, the image artifacts are of a type that may be accommodated in avariety of important imaging applications as they are displaced outsideof bright objects (with strong NMR signals). Accordingly, inapplications like angiography, where the bright vessels dominate thesignals from surrounding tissue, the image artifacts are acceptablebeing outside of the structures (i.e., the blood vessels) of interest.Artifacts from nearby tissues which are not as bright but would overlapthe blood vessels, are overwhelmed by the signal from the vesselsthemselves. Similar imaging applications are imaging of the ducts of thepancreas in a choloangiopancreatographic image or ducts of the liver.

Subtraction Studies

Referring now to FIGS. 12 and 16, the projection imaging of the presentinvention may be used in subtraction-type studies, such as those using acontrast medium injected into blood vessels. As is understood in theart, in such subtraction imaging, a “pre-contrast” image is firstobtained of the region without the injected contrast medium followed bya “contrast” image of the same region after the contrast medium hasinfused into the structure of interest. The two images are thensubtracted on a point by point basis to reveal an image principally ofthe contrast medium infused structure.

The process of subtraction tends to remove artifacts generated by spinsoutside of the region of interest 64 or by sparse sampling which remainthe same in the two images and thus provides further reduction of imageartifacts.

Referring now to FIG. 16, a contrast medium may be injected into thepatient at a time 116 and there may be a unpredictable delay before adesirable contrast level 114 is reached. Generally too, the amount oftime during which desirable contrast 114 is maintained may be less thanthe time required to obtain a full set of projections 82. In the presentinvention, a complete projection set may be acquired on anafter-the-fact basis as will now be described.

As is well understood in the art, a projection set capable ofreconstructing an image ideally includes a large number (e.g., 128) ofprojections 82 uniformly distributed in angle about the patient.

Referring to FIG. 17 in a preferred embodiment, the projection set isdivided into subsets of projections 82 according to angles of axis x′.Thus for example, if the projection set includes 128 projections 82acquired along x′ axes differing by substantially 2.8 degrees, theseprojections 82 may be divided into eight projection subsets p1 throughp8. Projection set p1 may include projections 82 at angles θ of 8N(2.8),where N is an integer from 0 to 15. Similarly, projection subset p2 mayinclude projections at angles θ of (8N+1)(2.8) and so forth.

Projection subsets p1 through p8 are acquired in sequence startingbefore the injection time 116 of the contrast media and continuing untila fall off of contrast 112. Each projection subset as so dividedprovides a good angular dispersion of projections 82 and on anafter-the-fact basis, a contrast set of projections 118 including ofeight contiguous projection subsets may be selected centered about theperiod of optimal contrast level as will form the contrast image. Notethat the contrast set of projections 118 need not begin with projectionsubset p1.

Alternatively, the projection subsets p1-p8 may simply be sets ofsequential adjacent projections 82, however, the good angular dispersionwill not then be obtained, such as may reduce motion and other types ofartifacts. Immediately prior to this contrast set of projections 118, asecond pre-contrast set of projections 120 may be collected representinga pre-contrast image of the patient.

Optionally each of these projection sets 120 and 118 may be weighted byweighting functions 122 which reduce the contribution of projectionsubsets as a function of their time of acquisition. For the contrast setof projections 118, the weighting is such as to reduce the contributionfrom projection subsets removed from the time of peak contrast 112. Forthe pre-contrast image projection subsets, the weighting may be such asto reduce the contribution from projection subsets beginning to exhibitadditional contrast as a result of infusion of contrast medium.

The two projection sets 120 and 118 as weighted may then be subtractedas indicated by block 124 and a back projection 86 made of thesubtracted image.

The contrast 112 may be monitored on a real-time basis by, afteracquiring the NMR data for each projection 82, summing or integratingthe kx′ data associated with that projection after the kx′ data has beentransformed. This sum may be compared to a sum for the identicalprojection (having the same 0 value) taken at a time prior to theinjection of the contrast medium. The difference between these valuesindicates the contrast index 112 with a time resolution equalsubstantially to the time required to acquire each new projection setafter the injection of the contrast medium. Analyses of the trending ofthe contrast 112 can be used to determined the appropriate collection ofthe projections into projection sets 120 and 118 and the appropriateweighting functions 122.

Alternatively a historical functional relationship of integrated kx′data to angle may be determined and used to weight the kx′ data so thatit may be compared to kx′ data at different angles to deduce changes incontrast and trending prior to the repeating of a projection set. Thiscontrast index value 112 may be displayed to the physician or used toselect sets of projections for projections sets to be reconstructed.

Volume Acquisitions

Images may be constructed of a volume extending in the z direction byacquiring successive slices in z-axis, back projecting slice images, andcombining several slices using Fourier transform techniques well knownin conventional imaging. Alternatively, the projection acquisition ofthe present invention may be combined with z-phase encoding todistinguish the z-dimension, according to techniques well understood inthe art. Two approaches may be used. In the first, a complete projectionset or projection subset is acquired for each z encoding. This permitsplacing all the low spatial frequency z encodes near the peak of thecontrast curve. In the second approach, the full range of z encodes isperformed for each projection of the projection set or projectionsubset. This method allows the after-the-fact collection of projectionsinto a projection set clustered about the peak of the contrast curve.Obviously combinations of these approaches may be used in whichz-encodes occur between subsets of projections.

Referring now to FIG. 18, in a third method, the amount of time requiredto acquire a volume of data encompassing several z-axis slices may bereduced by a k-space biased acquisition technique described in U.S. Pat.No. 5,713,358 issued Feb. 3, 1998 and entitled: METHOD FOR PRODUCING ATIME-RESOLVED SERIES OF 3D MAGNETIC RESONANCE ANGIOGRAMS DURING THEFIRST PASSAGE OF CONTRAST AGENT” naming a co-inventor of the presentapplication, assigned to the same assignee as the present invention, andhereby incorporated by reference. This technique involves the repetitiveacquisition of a volume 126 of data in k-space, for example, for thecontrast measurement described above with respect to FIG. 16. Thetechnique separately acquires k-space data in different z-axis bands, toobtain data in the kz-dimension near the origin 128 on a more frequentbasis than data obtained away from the origin.

Thus, the k-space data may be divided into three regions differing in kzlocation. “A-data” is that in a band closest to the origin of k-space,“B-data” is removed from the origin on either side of the A-data, and“C-data” is removed even further from the origin to flank the B andA-data. The acquisition of data over a volume through z then follows thepattern of ABACABAC, etc. so that A data are preferentially obtained.This acquisition reflects a recognition that low frequency image data isrelatively more important in contrast images then high frequency imagedata.

In acquiring data over a volume, the technique described with respect toFIG. 18 may be combined with the technique described with respect toFIG. 16 with each projection subset p being acquired over a volumepiece-wise in the k-z dimension.

It will be understood from this description that the present techniquemay be used with a number of different types of pulse sequences andimage acquisition protocols well known in the art but as applied to aradial acquisition as described. Thus, the technique may be used in asingle slice with successive radial acquisitions or may be used in avolume of excited spins by tilting the spokes out of the x-y plane intothe z-plane or may be used with z-phase encoding. Alternatively a volumeof data may be obtained with a series of successive single sliceexcitations. In the situation, for example, where the image plane isaligned with the x-y plane and the field of view is moved in thez-direction a helical acquisition of data may be obtained in a manneranalogous to that used in spiral or helical CT imaging. Thus, as motionis had in the z-direction, successive spokes may be obtained in ahelical pattern with the 180° of spokes closest to the desired imageplane being used to reconstruct the image or new spokes replacing oldspokes as additional z-direction is reached. The spokes not lying in theimage plane may be interpolated or extrapolated to the image plane asdesired.

Further it will be understood that the above techniques associated withreduced field of view imaging and full field of view imaging maybeinterchanged and that no element of these techniques not claimed shouldbe considered critical to the invention.

The above description has been that of a preferred embodiment of thepresent invention. It will occur to those that practice the art thatmany modifications may be made without departing from the spirit andscope of the invention. In order to apprise the public of the variousembodiments that may fall within the scope of the invention, thefollowing claims are made.

We claim:
 1. A method for producing a magnetic resonance angiogram ofselected vasculature in a subject using a magnetic resonance imaging(MRI) system, the steps comprising: a) acquiring a mask image of thevasculature using the MRI system; b) introducing a contrast agent whichflows into the selected vasculature; c) operating the MRI system toperform a pulse sequence which includes: i) producing an RF excitationpulse to excite spins in a field of view which includes the selectedvasculature; ii) applying a phase encoding gradient along a first axis;iii) applying a radial gradient directed at an angle θ in a planeperpendicular to the first axis; and iv) acquiring an NMR signal duringthe application of the radial gradient to sample an angular projectionof the data in k-space having data points radially spaced at distance NRalong a projection; d) repeating step c) with a set of different phaseencoding gradient values for each of a plurality of different radialgradient angles θ until k-space is sampled, wherein the plurality ofdifferent radial gradient angles θ is less than NR π/4 in number so thata sparsely sampled k-space data set is acquired; e) reconstructing animage of the selected vasculature from the sparsely sampled k-space dataset; and f) subtracting the mask image from the image reconstructed instep e).
 2. The method as recited in claim 1 in which the pulse sequenceused to operate the MRI system in step c) is also used to acquire themask image in step a).
 3. The method as recited in claim 1 in which stepd) is repeated to acquire a corresponding plurality of sparsely sampledk-space data sets during the passage of the contrast agent through theselected vasculature; and step e) includes combining k-space data from aplurality of said sparsely sampled k-space data sets.
 4. The method asrecited in claim 3 in which the phase encoding gradient values forencoding lower k-space frequencies are applied disproportionately morefrequently than the phase encoding values for higher k-space frequenciesduring the acquisition of said plurality of sparsely sampled k-spacedata sets.
 5. The method as recited in claim 3 in which the phaseencoding gradient values for encoding for higher k-space frequencies aredifferent in successive ones of said plurality of sparsely sampledk-space data sets.
 6. A method for producing a magnetic resonanceangiogram of selected vasculature in a subject using a magneticresonance imaging (MRI) system, the steps comprising: a) introducing acontrast agent which flows into the selected vasculature; b) operatingthe MRI system to perform a pulse sequence which includes: i) producingan RF excitation pulse to excite spins in a field of view which includesthe selected vasculature; ii) applying a phase encoding gradient along afirst axis; iii) applying a radial gradient directed at an angle θ in aplane perpendicular to the first axis; and iv) acquiring an NMR signalduring the application of the radial gradient to sample an angularprojection of the data in k-space having data points radially spaced atdistance NR along a projection; c) repeating step b) with a set ofdifferent phase encoding gradient values for each of a plurality ofdifferent radial gradient angles θ until k-space is sampled, wherein theplurality of different radial gradient angles θ is less than NR π/4 innumber so that a sparsely sampled k-space data set is acquired; d)repeating step c) a plurality of times to acquire a correspondingplurality of sparsely sampled k-space data sets during the passage ofthe contrast agent through the selected vasculature, and whereinsuccessive sparsely sampled k-space data sets sample different parts ofk-space; e) combining k-space data from a plurality of said sparselysampled k-space data sets acquired during a selected time intervalduring the passage of the contrast agent through the selectedvasculature; and f) reconstructing an image of the selected vasculatureusing the combined k-space data from step e).
 7. The method as recitedin claim 6 in which the phase encoding gradient values for encoding forhigher k-space frequencies are different in successive ones of saidplurality of sparsely sampled k-space data sets.
 8. The method asrecited in claim 7 in which the radial gradient angles θ produced duringthe acquisition of each sparsely sampled k-space data set are differentfor successively acquired sparsely sampled k-space data sets.
 9. Themethod as recited in claim 6 in which step e) includes weighting thek-space data differently for each of the plurality of sparsely sampledk-space data sets.